Method, system and software arrangement, for measuring magnetic field correlation

ABSTRACT

Methods, systems, software arrangements and storage medium for measuring the magnetic field correlation function (“MFC”), and more particularly, to methods for measuring the magnetic field correlation function utilizing asymmetric spin echoes. Asymmetric Dual Spin Echo Sequences (“ADSE”) and Echo Planar Imaging Asymmetric Dual Spin Echo Sequences (“EPI-ADSE”) may be employed to apply multiple echoes to a sample and acquire data from which the MFC may be determined.

CROSS-REFERENCE TO RELATED APPLICATION

The present application also claims priority from U.S. PatentApplication No. 60/485,502, filed Jul. 8, 2003, the entire disclosure ofwhich incorporated herein by reference.

FIELD OF THE INVENTION

The present invention relates to methods, systems, software arrangementsand storage medium for measuring a magnetic field correlation functionin a magnetic resonance imaging area, and more particularly, to methodsfor measuring the magnetic field correlation (“MFC”) function utilizingasymmetric spin echoes.

BACKGROUND OF THE INVENTION

A theory of using asymmetric spin echoes is described, in part, inJensen J. H. et al. Chandra R. Magn Reson Med; 44:144 (2000) (the“Jensen Publication”), the entire disclosure of which is incorporatedherein by reference.

The local magnetic field experienced by a water proton is quantitativelycharacterized by the magnetic field correlation function, K(t), whichmay be defined byK(|t−t′|)=

B(t)B(t′)

,  [1]with B(t) being the difference, at a time t, between the magnitude ofthe local field and the magnitude of the spatially uniform main field,B₀, and with the angle brackets indicating an averaging over all thewater protons within a given region of interest. K(t) may depend on boththe spatial distribution of the magnetic field inhomogeneities and thediffusional dynamics of water molecules. The magnetic field correlationfunction provides information beyond the information contained in thestandard nuclear magnetic resonance (“NMR”) relaxation times. Thus, thepresent invention provides a novel technique to, inter alia, examine theproperties of tissue, blood, iron-rich regions of the brain, and tumors.

Although it is believed that the MFC has not been specifically measured,it was introduced at least as early as 1953 in the seminal work ofAnderson and Weiss on NMR line shapes. See Anderson P W, Weiss P R, RevMod Phys 1953; I25:269-276 (the “Anderson Publication”), the entiredisclosure of which is incorporated herein by reference. More recently,it has been utilized in several studies on the modeling of MRI contrast.For example, see Callaghan P T, Oxford University Press, New York, 1991;(the “Callaghan Publication”) Kennan et al., Magn Reson Med 1994;I31:9-21 (the “Kennan Publication”); Stables et al., Magn Reson Med1998; I40:432-442 (the “Stables Publication”); and the JensenPublication et al., the entire disclosures of which are incorporatedherein by reference.

The MFC may be described as the magnetic resonance (“MR”) signalintensity as a function of the acquisition time, and can be approximatedby the exponential form:K(t)=K ₀exp(−t/τ),  [2]where K₀=K(0) is the magnetic field variance and τ is a characteristicdecay time. It has recently been shown in the Jensen Publication, thatwhen water diffusion is only weakly restricted the MFC is moreaccurately described by an algebraic expression of the form:$\begin{matrix}{{K(t)} = {{K_{0}\left( {1 + \frac{t}{\tau}} \right)}^{{- 3}/2}.}} & \lbrack 3\rbrack\end{matrix}$There is not a significant amount of quantitative information regardingthe MFC, except for an exact result that can be derived for an idealizedrandom sphere model.

Asymmetric single spin echoes were first introduced by Dixon andSepponen. In particular, Dixon demonstrated how asymmetric spin echoescan be used to separate the MR signals originating from water and fat.See Dixon W T, Radiology 1984; I153:189-194 (the “Dixon Publication”),the entire disclosure of which is incorporated herein by reference.Sepponen et al., applied asymmetric spin echoes to obtain chemical shiftimages. See Sepponen R E, et al., Comput. Assist. Tomography 1984;I8:585-587 (the “Sepponen Publication”), the entire disclosure of whichis incorporated herein by reference. The asymmetry of the Dixon sequencearises by shifting the signal acquisition time, while the asymmetry ofthe sequence of Sepponen et al. is achieved by shifting the 180°refocusing pulse.

Asymmetric single spin echo technique with a shifted refocusing pulsehas been previously used to obtain MFC data. See Wismer et al., J ComputAssist Tomography 1988; I12:259-265 (the “Wismer Publication”);Rosenthal et al., Invest Radiology 1990; I025:173-178 (the “RosenthalPublication”); Thulborn et al., Am J Neuroradiol. 1990 (the “ThulbornPublication”); 11:291-297; Hoppel et al., Magn Reson Med 1993;I30:715-723 (the “Hoppel Publication”); Ganesan et al., J. Magn. Reson.(B) 1993; I102:293-298 (the “Ganesan Publication”); and the StablesPublication, the entire disclosures of which are incorporated herein byreference. However, there remains a need for improved methods ofmeasuring the magnetic field correlation.

SUMMARY OF THE INVENTION

The present invention relates to methods, systems, software arrangementsand storage medium for measuring the magnetic field correlationfunction, and more particularly, to methods, systems, softwarearrangements and storage medium measuring the magnetic field correlationfunction utilizing asymmetric spin echoes.

In an exemplary embodiment of the present invention, method, systems,software arrangements and storage medium obtaining the magnetic fieldcorrelation (“MFC”) of a sample using magnetic resonance for imaging(“MRI”) are provided that apply multiple spin echo sequences in which atleast one spin echo sequence is an asymmetric spin echo sequence.Resultant information may then be acquired and the MFC can be determinedas a function of the acquired information.

In a further exemplary embodiment of the present invention, the spinecho sequences may be an Asymmetric Dual Spin Echo Sequence (“ADSE”)having multiple echoes. In another exemplary embodiment of the presentinvention, the spin echo sequences may be an Echo PlanarImaging-Asymmetric Dual Spin Echo Sequence (“EPI-ADSE”) having multipleechoes.

In a further embodiment of the present invention, the asymmetric spinecho sequence can be applied by shifting a refocusing pulse such thatthe time between the rotation pulse and the refocusing pulse (t₁) is notequal to the time between the refocusing pulse and the signalacquisition (t₂), e.g., t₁≠t₂. In yet another exemplary embodiment, theasymmetric spin echo sequence may be applied by shifting the acquisitionof the resultant information such that the time between the rotationpulse and the refocusing pulse (t₁) is not equal to the time between therefocusing pulse and the signal acquisition (t₂), i.e., t₁≠t₂.

In a further exemplary embodiment, the MFC can be determined as afunction of the resultant information by applying the formula${{K\left\lbrack {\left( {{2n} - 1} \right)\Delta\quad t} \right\rbrack} \approx {\frac{\left( {- 1} \right)^{n + 1}}{2\gamma^{2}t_{s}^{2}}{\ln\left\lbrack \frac{{S_{n}(0)}{S_{n - 1}\left( t_{s} \right)}}{{S_{n}\left( t_{s} \right)}{S_{n - 1}(0)}} \right\rbrack}}},$in which γ is the proton gyromagnetic ratio, S_(n) is the signalintensity of the nth echo; and t_(s)=|t₁−t₂| with t₁ being the timebetween a rotation pulse and a refocusing pulse, and t₂ being the timebetween the refocusing pulse and a signal acquisition.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1(a) shows an exemplary illustration of randomly distributedspheres representing cells with a higher magnetic susceptibility thenthe surrounding tissue;

FIG. 1(b) shows exemplary contour lines for the magnetic field shiftsinduced by an external field oriented in the z direction, with solidlines provided for positive shifts, and dotted lines provided fornegative shifts;

FIG. 1(c) shows diffusion paths for two water molecules beginning at atime t and ending at a time t′;

FIG. 1(d) shows an illustration of an exemplary magnetic fieldcorrelation (“MFC”) as a function of time representing a decay of thefield shift correlation;

FIG. 2 shows an exemplary asymmetric Carr-Purrcell-Meiboom-Gill (CPMG)sequence with N=3;

FIG. 3(a) shows an exemplary asymmetric spin echo signal intensityprovided as a function of the refocusing pulse time shift for tworegions of interest within a subacute hemorragic brain lesion; and

FIG. 3(b) shows an exemplary semi-logarithmic graph provided as afunction of the square of the time shift

FIG. 4 shows a flow diagram representing an exemplary method andprocedure under at least partial control of a computing arrangement ofFIG. 5 using the methods and systems of the present invention;

FIG. 5 shows a schematic block diagram of an exemplary embodiment of asystem according to the present invention.

Other and further objects, features and advantages of the presentinvention will be readily apparent to those skilled in the art upon areading of the description of preferred embodiments which follows.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention relates to methods, systems, software arrangementsand storage medium for measuring the MFC function, and moreparticularly, to methods, software arrangements and storage medium formeasuring the MFC function utilizing asymmetric spin echoes.

Since the MFC is sensitive to the spatial distribution of microscopicmagnetic field inhomogeneities, MFC imaging enables another way to probetissue microstructures. For example, MFC imaging may be potentiallyuseful for evaluating tissues with substantial intrinsic microscopicmagnetic susceptibility variations. It may therefore be used as asensitive probe of tissue microstructure that provides certaininformation beyond what is contained in the standard nuclear magneticresonance (“NMR”) relaxation times, such as T2 and T2*. In addition,microscopic susceptibility variations can be created by the introductionof a contrast agent. For example, the contrast agent may be aparamagnetic element, e.g., gadopentetate dimeglumine (Gd-DTPA).Therefore, the imaging of the MFC may be useful in the evaluation ofneurodegenerative disorders, such as Parkinson's and Alzheimer'sdiseases, since ferritin (i.e., the most prevalent form of iron in thebrain) is paramagnetic and localized in oligodendrocytes, which havediameters of about 7 μm. Indeed, it is known that the T2 relaxationrate, as measured with a Carr-Purrcell-Meiboom-Gill (CPMG) sequence,depends significantly on the interecho time in the brain region. Thus,the MFC with a nontrivial time dependence may be expected and themeasurement of this time dependent MFC can provide a significant markerof disease state.

The MFC imaging may be of particular interest for Alzheimer's disease,as conventional relaxation time measurements may not be as useful.Another possibly important application the MFC imaging may be to theimaging of tumors. Most paramagnetic contrast agents, such as Gd-DTPA,enter only the extra cellular spaces of tissues, and thereby createmicroscopic magnetic susceptibility variations. The MFC of a contrastagent infused tumor may be sensitive to tumor morphology (e.g., cellsize and density). The measurement of the MFC can therefore assist intumor identification and grading. Moreover, measuring the MFC with themagnetic resonance imaging (“MRI”) may be a preferable way for probinginhomogeneities on a length scale of a few tens of micrometers. This maybe because MRI echo times are typically in the range of 10 ms to 100 ms,and the water diffusion constant in biological tissues is about 1.0μm²/ms. This implies that the temporal behavior of the measured MFC maybe sensitive to spatial variations in the range of 8 μm to 25 μm.Biologically, this may be an important length scale, since this is thesize of capillaries and many types of cells.

The MFC may be defined as the product of the magnetic field shiftsexperienced by a water molecule at two different times, averaged overall the water molecules within a specified region of interest. The fieldshifts may be the differences between the magnitude of the field and themagnitude of the uniform background field (which may be nearly the sameas the external field magnitude). If the biological tissue has linearmagnetic characteristics, as is usually the case, then the MFC may varyas the square of the external field. The dependence of the MFC on thedifference in the two sampling times, on the other hand, may be purelyan intrinsic tissue property. In practice, the MFC may be most sensitiveto the spatial dependence of the field shifts over length scales ofabout 10 μm, which is comparable to the size of many cell types.

To appreciate the significance of the MFC, it is preferable tounderstand the definition of the MFC, which is provided as follows. Forexample, a tissue may be exposed to a strong, uniform external magneticfield, as is the case in a typical MRI experiment. Because of spatialvariations in its magnetic susceptibility, the tissue may generate aninhomogeneous secondary field. The susceptibility variations can beprovided due to structures such as iron-rich cells or vessels with adeoxygenated blood. A water molecule diffusing through the tissue mayexperience a field that varies with time due to field inhomogeneities.If δB(t) is defined as the difference at time t between the magnitude ofthe field at the water molecule and the uniform background field, thenthe MFC may be defined byK(|t−t′|)=

δB(t)δB(t′)

,  [4]where the angle brackets indicate an average over all the watermolecules in a specified region of interest (in practice this may be anMRI voxel). The temporal behavior of the MFC, which is indicated with asymbol K in mathematical expressions, involves the time difference|t−t′|, as follows from an assumption that the state of the tissue maybe time invariant. It should be noted that, by definition, the MFC candepend quadratically on the magnitude of the field inhomogeneities.

As an illustration, FIG. 1(a) depicts an exemplary set of approximatelyidentical spheres distributed randomly in space. The random spheres mayserve, for example, as models for iron-rich glial cells as described inJensen J H, et al., Magn Reson Med 2001; I46:159-165 (the “Second JensenPublication”), the entire disclosure of which is incorporated herein byreference. For example, the magnetic susceptibility of the spheres maybe higher than that of the surrounding media, and a uniform magneticfield is applied in the z direction. An inhomogeneous secondary magneticfield (FIG. 1(b)) can then be generated. FIG. 1(c) depicts an exemplarydiffusion path of two water molecules. For example, diffusion paths fortwo water molecules beginning at a time t and ending at a time t′. When|t−t′| is small, the diffusion path is short, and the initial and finalfield shifts are strongly correlated. When |t−t′| is large the diffusionpath is long and the initial and final field shifts are weeklycorrelated. For the shorter path, which corresponds to a smaller timedifference between the initial and final points, the initial and finalmagnetic field shifts are similar, indicating a high degree ofcorrelation. For the longer path, the magnetic field shifts aredifferent, indicating a low degree of correlation. FIG. 1(d) plots thecorresponding decay of the MFC as a function of time.

According to one exemplary embodiment according to the presentinvention, the MFC at specific times can be directly imaged withoutrecourse to any particular model. This exemplary technique thereby yieldinformation beyond what may be provided by conventional relaxation timemeasurements. The fact that the MFC contains information distinct fromT2 and T2* can be described using the following simple example. Forexample, two test tubes, A and B, can be provided each containing anaqueous suspension of microspheres and equal concentrations of aparamagnetic contrast agent. If the microspheres in test tube A have aradius of 10 μm and the microspheres in test tube B have a radius of 20μm, the spatial density of spheres in test tube A can be eight timeshigher than in test tube B. The volume fraction filled by the spherescan then be approximately the same in both suspensions. T2 can be thesame for both test tubes because T2 can be determined primarily by thecontrast agent concentration. T2* can also be essentially the same,because T2* may depend primarily on the contrast agent concentration andthe volume fraction of the spheres. Therefore, the test tubes may appearapproximately identically on T2 or T2* weighted images. However, if awater diffusion constant of 1.0 μm²/ms is assumed, then for a timedifference of 20 ms the MFC values for the two test tubes will differ byabout a factor of two, and the test tubes may appear sharply differenton an MFC image.

A relationship between the MFC and transverse relaxation rates may bemost clearly revealed through a connection between the MFC and T2 asmeasured with a Carr-Purcell-Meiboom-Gill (CPMG) sequence. Because ofwater diffusion through field gradients, the measured T2, as obtainedfrom a monoexponential fit to the echo intensities, may depend on theinterecho time, 2Δt. In the limit 2Δt→0, the true T2 value may beobtained. If the magnetic field inhomogeneities are not too large, thenone can show that T2 as a function of Δt may be provided by$\begin{matrix}{{\frac{1}{T\quad 2\left( {\Delta\quad t} \right)} = {\frac{1}{T\quad 2(0)} + {\frac{8\gamma^{2}}{\pi^{2}}{\sum\limits_{m = 0}^{\infty}{\frac{1}{\left( {{2m} + 1} \right)^{2}}{\int_{0}^{\infty}\quad{{\mathbb{d}{{tK}(t)}}{\cos\left\lbrack \frac{\left( {{2m} + 1} \right){\pi t}}{2\Delta\quad t} \right\rbrack}}}}}}}},} & \lbrack 5\rbrack\end{matrix}$with γ representing the proton gyromagnetic ratio.

If the full dependence of T2 on the interecho time were available, thenit is possible to invert Eq. [5] to determine the MFC. However, it isgenerally only practical to obtain T2 for a limited number of Δt values,and so a straightforward inversion may not be possible. In some cases,it may be possible to postulate a model form for the MFC, having a fewadjustable parameters, and then fit the model to the T2 data by usingequation [5]. However, the success of this approach may depend on theaccuracy of the model, which can be difficult to establish.

The physical significance of the MFC can be twofold. First, themagnitude of the MFC may provide a measure of the amplitude of themagnetic field inhomogeneities within a tissue. In particular, asfollows from Eq. [3], K(0) can be the variance of the field. Second, thedecay of the MFC with time may contain information about the spatialvariation of the field inhomogeneities. In particular, the temporal rateof change of the MFC at a time t (i.e., K′(t)) can be primarilysensitive to field inhomogeneities with a length scale of about √{squareroot over (6Dt)}, where D is the water diffusion constant, where√{square root over (6Dt)} being the average distance a water moleculediffuses over a time t. Thus, the MFC may be measured in conjunctionwith diffusion-weighted imaging (DWI) in order to obtain an estimate ofthe water diffusion constant. This additional information may then allowfor a more accurate assessment of the length scale associated with thefield inhomogeneities. If the inhomogeneities vary only on length scalesthat are large compared to about √{square root over (6Dt)} (i.e., >>25μm), the MFC can be essentially independent of time and may be relatedto the NMR line width. Particularly, the second moment of the NMRspectral line may be approximately about 4γ²K.

According to an exemplary embodiment of the present invention, the MFCmay be determined by applying 2 or more spin echo sequences, in which atleast one spin echo sequence is an asymmetric spin echo sequence. FIG. 2shows an exemplary asymmetric Carr-Purrcell-Meiboom-Gill (CPMG) sequencewith N=3. The 180° refocusing pulses are shifted by a time t_(s) whilethe acquisition time is unchanged. According to one exemplary embodimentof the present invention, a spin echo sequence can be provided in whichan initial 90° rotation pulse may be followed at a time t_(p) by a 180°refocusing pulse and an NMR signal of magnitude S(t, t_(p)) is collectedat a time t. For standard spin echo imaging, t can be selected to beequal to 2t_(p). Thus, an asymmetric spin echo image may have aboutt≠2t_(p) and the following formula may be applied: $\begin{matrix}{{{In}\left\lbrack \frac{S\left( {{2t},t} \right)}{S\left( {{2t},{t + t_{s}}} \right)} \right\rbrack} \approx {2\gamma^{2}{\int_{t_{s}}^{t_{s}}\quad{\mathbb{d}{t^{\prime}\left( {{t_{s} - {{t^{\prime}}{K\left( {t + t^{\prime}} \right)}}},} \right.}}}}} & \lbrack 6\rbrack\end{matrix}$where γ is the proton gyromagnetic ratio.

Equation [6] is based on a weak field approximation, but is also usableat clinical field levels of a few Tesla or less. The correlationfunction enters equation [6] preferably only for times between t−ts andt+ts. If K(t) is approximated in this temporal range by a linearfunction, the following formula may be applied: $\begin{matrix}{{\frac{1}{2\gamma^{2}t_{s}^{2}}1{n\left\lbrack \frac{S\left( {{2t},t} \right)}{S\left( {{2t},{t + t_{s}}} \right)} \right\rbrack}} \approx {{K(t)}.}} & \lbrack 7\rbrack\end{matrix}$Hence, the ratio of the signal intensities for a standard spin echosequence and for an asymmetric spin echo sequence can be used toestimate K(t) at a selected time. This exemplary technique may beutilized for multiple spin echo sequences so that K(t) at severaldifferent times may be determined from one pair of acquisitions.

For example, the key may be a relationship between the MFC and thesignal intensity for an asymmetric CPMG sequence. The asymmetric CPMGsequence differs from a conventional CPMG sequence in that the 180°refocusing pulses are shifted from their usual times by an amount t_(s),as is illustrated in FIG. 2. Regardless of t_(s), the signal can besampled at times 2nΔt, for n=1, 2, . . . , N, where 2Δt is the interechotime and N is the total number of echoes. Now let S_(n)(t_(s)) be themagnitude of the nth echo. $\begin{matrix}{{{\int_{- t_{s}}^{t_{s}}\quad{{\mathbb{d}{t\left( {t_{s} - {t}} \right)}}{K\left\lbrack {\left( {t + {2n} - 1} \right)\Delta\quad t} \right\rbrack}}} \approx {\frac{\left( {- 1} \right)^{n + 1}}{2\gamma^{2}}{\ln\left\lbrack \frac{{S_{n}(0)}{S_{n - 1}\left( t_{s} \right)}}{{S_{n}\left( t_{s} \right)}{S_{n - 1}(0)}} \right\rbrack}}},} & \lbrack 8\rbrack\end{matrix}$with the convention S⁻¹(ts)=1. The corrections to equation [8] arepreferably third order in the field, and for weak fields are smallcompared to the MFC, which is second order in the field. The integral inequation [8] is over the time t ranging from −t_(s) to t_(s). As long ast_(s) is not too large, it is reasonable to make the following linearapproximation:K[(t+2n−1)Δt]≈K[(2n−1)Δt]+tK′[(2n−1)Δt],  [9]which leads to the following simplification: $\begin{matrix}{{{K\left\lbrack {\left( {{2n} - 1} \right)\Delta\quad t} \right\rbrack} \approx {\frac{\left( {- 1} \right)^{n + 1}}{2\gamma^{2}t_{s}^{2}}{\ln\left\lbrack \frac{{S_{n}(0)}{S_{n - 1}\left( t_{s} \right)}}{{S_{n}\left( t_{s} \right)}{S_{n - 1}(0)}} \right\rbrack}}},} & \lbrack 10\rbrack\end{matrix}$

This expression indicates that the MFC at the time of the nth refocusingpulse can be estimated from the signal intensities of the preceding andsucceeding echoes. As provided in equation [10], an echo train of Nechoes can be used to obtain the MFC at N different times.

As an example of how to apply equation [10] to obtain an estimate forthe MFC, the experiment described in Wismer may be considered. SeeWismer et al., J Comput Assist Tomography 1988; I12:259-265, the entiredisclosure of which is incorporated herein by reference. FIG. 3 a showsthe signal intensity from a subacute hemorrhagic brain lesion obtainedin vivo with asymmetric spin echoes. The field level was 0.6 T, the TEmay be 50 ms, and the TR may be 500 ms. For n=1, Eq. [10] can beexpressed as1n[S ₁(t _(s))]≈1n[S ₁(0)]−2γ² t _(s) ² K(Δt).  [11]Thus, the signal intensity is preferably a linear function of t_(s) ² ona semi-logarithmic plot. An estimate of the MFC at the time of therefocusing pulse can then be found from the slope of a line fitted tothe data. In FIG. 3 b, the same data shown in FIG. 3 a is replotted as afunction of t_(s) ². The slope of a linear fit to the data allowing anestimate of the MFC at a time of 25 ms. From the slope of the linearfit, it is indicated that γ²K(25 ms)=6100±200s⁻², for the first regionof interest, and γ²K(25 ms)=3400±100s⁻², for the second region ofinterest. (When giving quantitative values for the MFC, a factor of γ²is included for convenience.)

Qualitatively similar behavior for the signal intensity as a function oft_(s) has been reported for normal brain, bone marrow, lung tissue andmicrosphere phantoms. See the Hoppel, Rosenthal, Ganesan., and StablesPublications.

Referring to FIG. 2, the RF (i.e., radio frequency) pulses, X magneticgradient, Y magnetic gradient, and Z magnetic gradient are plottedagainst time. An initial magnetic field polarizes the molecules in the Zdirection. A 90 degree rotation pulse can then be used to move themagnetization into the XY plane, which precesses about the Z direction.A 180 degree refocusing pulse may then be used to correct a phasedispersion. Additional refocusing pulses may also be used to increasethe signal, and further correct the phase dispersion. In an exemplaryembodiment of the present invention, at least two refocusing pulses maybe used. In a further embodiment, the gradient coils can be energized atvarious times to produce an image of the subject being analyzed.

One of the features of the present invention is the use of the RF pulsesequence. Referring to FIG. 2, the dotted line shown therein representsthe typical time that a refocusing pulse is used in conventionalsystems. This is approximately provided between the time of the rotationpulse and the time the signal is acquired. In an exemplary aspect of thepresent invention, the radio frequency pulse can be shifted by timet_(s). The imaging signal changes as the pulse is shifted eitherpositively or negatively. In a further exemplary embodiment, therefocusing pulse may be shifted in one direction up to the point whereit is approximately simultaneous with the 90 degree pulse. In anotherembodiment of the present invention the refocusing pulse may be shiftedby an amount up to the point where it is nearly simultaneous with thesignal acquisition. In yet another exemplary embodiment of the presentinvention, the refocusing pulse may be shifted back and fortharbitrarily within that region. In still another embodiment of thepresent invention, the refocusing pulse may be shifted about ¼ of thetime between the previous pulse and the time the signal is obtained.

According to one exemplary embodiment of the present invention, twopulse sequence programs implementing an ADSE may be used. A dual spinecho has the advantage over a single spin echo by simultaneouslyproviding the information usable to estimate the MFC for two differenttime values. In one exemplary embodiment of the present invention, aconventional sequence with a single phase encoding step for each 90°radio frequency (RF) pulse may be used. For example, the sequence shownin FIG. 2 may be utilized, with only two 180° pulses. In anotherexemplary embodiment of the present invention, an EPI dual spin echo inwhich all of phase space can be acquired with a single 90° pulse. Sincethe image acquisition time is a fraction of a second for the EPIsequence, as compared to several minutes for the conventional sequence,the EPI sequence may be more convenient for some applications. Inaddition, the EPI sequence may reduce associated motion artifacts.

Using the exemplary techniques described in the present invention, theMFC may be estimated by applying equation [7] to the ADSE imaging data.For example, typical MFC values obtained in vivo for a human brain areusing the methods and systems according to the present invention listedin Table 1. One prior in vivo human study, which allows for MFCestimates, gives data just for hemorrhagic brain lesions, which have amuch higher MFC than healthy brain tissue. See Wismer Publication. Thisstudy, however, has been performed at a low field level of 0.6 T, andthe MFC in normal brain was presumably too small to observe due to asmall SNR. Since the MFC increases quadratically with the fieldstrength, the MFC at 1.5 T is 6.25 times larger than at 0.6 T, itsdetection is easier. TABLE 1 Brain Region Mean MFC (1/s²) GlobusPallidus 1004 ± 112 Substania Nigra  820 ± 92 Red Nucleus  715 ± 75Putamen  497 ± 83 Caudate Head  443 ± 73 Thalamus  318 ± 100 CorticalGray Matter  268 ± 95 Frontal White Matter  150 ± 104

FIG. 4 is a flow diagram representing an exemplary procedure under atleast partial computer control using the methods and systems of thepresent invention, as may be carried out by the system of FIG. 5. Instep 410, a spin echo sequence is applied to the sample. Next, thecorresponding resultant information is acquired in step 420. In step430, another spin echo sequence is applied to the sample. Then, in step440 resultant information is acquired corresponding to the spin echosequence applied in step 430. In step 450, it is determined whether itis desirable to apply more spin echo sequences to the sample. If yes,additional spin echo sequences are applied and corresponding resultantinformation is further acquired. If no, the magnetic field correlationis determined based on the acquired resultant information. It should benoted that according to an exemplary embodiment of the present inventionat least one spin echo sequence is an asymmetric spin echo sequence.

FIG. 5 shows a schematic block diagram of an exemplary embodiment of asystem according to the present invention. In particular, an MRIinstrument 500 is controlled by a processing arrangement 510 and may beconnected to an output display 520. The processing arrangement maycontrol the spin-echo sequences applied to the sample located in the MRIinstrument 500 and also obtain the resultant information from thesample. The processing arrangement 510 may also be used to calculate themagnetic field correlation. In a further embodiment, the processingarrangement 510 may be used to generate an image as a function of thedetermined MFC, which may then be displayed on output display 520.

Therefore, the exemplary embodiment of the system process and softwarearrangement according to the present invention is well-adapted to carryout the objects and attain the ends and advantages mentioned as well asthose which are inherent therein. While the invention has been depicted,described, and is defined by reference to exemplary embodiments of theinvention, such a reference does not imply a limitation on theinvention, and no such limitation is to be inferred. The invention iscapable of considerable modification, alteration, and equivalents inform and function, as will occur to those ordinarily skilled in thepertinent arts and having the benefit of this disclosure. The depictedand described embodiments of the invention are exemplary only, and arenot exhaustive of the scope of the invention. Consequently, theinvention is intended to be limited only by the spirit and scope of theappended claims, giving full cognizance to equivalence in all respects.

1. A method for obtaining a magnetic field correlation (“MFC”) of asample using magnetic resonance imaging (“MRI”) comprising: applying twoor more spin echo sequences to the sample to obtain a resultantinformation, wherein at least one spin echo sequence is an asymmetricspin echo sequence; and determining the MFC as a function of theresultant information.
 2. The method of claim 1, wherein the spin echosequences include an Asymmetric Dual Spin Echo Sequence (ADSE) havingmultiple echoes.
 3. The method of claim 1, wherein the spin echosequences include an Echo Planar Imaging-Asymmetric Dual Spin EchoSequence (EPI-ADSE) having multiple echoes.
 4. The method of claim 1,wherein the asymmetric spin echo sequence is applied by shifting arefocusing pulse that is applied to the sample wherein a first timebetween a rotation pulse that is applied to the sample and therefocusing pulse is not equal to a second time between the refocusingpulse and obtaining the resultant information.
 5. The method of claim 1,wherein the asymmetric spin echo sequence is applied by shiftingobtaining of the resultant information wherein a first time between arotation pulse that is applied to the sample and the refocusing pulse isnot equal to a second time between the refocusing pulse and obtainingthe resultant information.
 6. The method of claim 1, wherein the MFC isdetermined as a function of the resultant information by applying theformula $\begin{matrix}{{{K\left\lbrack {\left( {{2n} - 1} \right)\Delta\quad t} \right\rbrack} \approx {\frac{\left( {- 1} \right)^{n + 1}}{2\gamma^{2}t_{s}^{2}}{\ln\left\lbrack \frac{{S_{n}(0)}{S_{n - 1}\left( t_{s} \right)}}{{S_{n}\left( t_{s} \right)}{S_{n - 1}(0)}} \right\rbrack}}},} & \lbrack 10\rbrack\end{matrix}$ wherein γ is the proton gyromagnetic ratio, S_(n) is thesignal intensity of the nth echo; t_(s)=|t₁−t₂|, where t₁ is the timebetween a rotation pulse that is applied to the sample and a refocusingpulse that is applied to the sample and t₂ is the time between therefocusing pulse and obtaining the resultant information.
 7. The methodof claim 1, further comprising generating an image as a function of thedetermined MFC.
 8. The method of claim 1, further comprising determininga distribution of a paramagnetic element in the sample as a function ofthe determined MFC.
 9. The method of claim 1, further comprisingdetermining a distribution of iron in the sample as a function of thedetermined MFC.
 10. The method of claim 1, further comprising adding acontrast agent to the sample prior to applying the spin echo sequences.11. The method of claim 10, wherein the contrast agent is gadopentetatedimeglumine (“Gd-DTPA”).
 12. The method of claim 1, further comprisingclassifying a tumor in the sample.
 13. A system for obtaining a magneticfield correlation (“MFC”) of a sample using magnetic resonance imaging(“MRI”) comprising: a storage medium, wherein the storage mediumincludes software that is capable of being executed to perform stepscomprising: applying two or more spin echo sequences to the sample toobtain a resultant information, wherein at least one spin echo sequenceis an asymmetric spin echo sequence; and determining the MFC as afunction of the resultant information. 14-24. (canceled)
 25. A softwarearrangement which, when executed on a processing device, configures theprocessing device to measure a magnetic field correlation (“MFC”) of asample using magnetic resonance imaging (“MRI”) comprising a set ofinstructions which when executed by the processing device perform stepscomprising: applying two or more spin echo sequences to the sample toobtain a resultant information, wherein at least one spin echo sequenceis an asymmetric spin echo sequence; and determining the MFC as afunction of the resultant information. 26-36. (canceled)
 37. A methodfor obtaining a magnetic field correlation (“MFC”) of a sample,comprising: applying two or more magnetic resonance imaging sequences toa predetermined region of the sample at a plurality of points in time toproduce resultant data; and determining the MFC as a function of atleast one set of molecules provided in the sample and the resultantdata.
 38. The method according to claim 37, wherein the moleculesinclude at least one of water molecules or fluorine molecules.
 39. Themethod according to claim 37, wherein the magnetic resonance imagingsequences include spin echo sequences.